1. Field of the Invention
The invention relates to a method for a computed tomography apparatus with a universal computer for the post-processing of a reconstructed tomogram of a slice of a subject of examination, which represents the entire field of measurement of the computed tomograph apparatus or a segment of the measurement field, based on an adaptive ring suppression filter.
2. Description of the Prior Art
Computed tomography is a standard technique in medicine, and is of great practical importance, e.g. for diagnosis. With a radiological measurement system including an X-ray detector system, which generally moves continuously around a subject of examination on a positioning table, or about the rotational center of the computed tomography apparatus, attenuation values of slices of the subject of examination are recorded at various angular positions of the measurement system relative to the subject. These attenuation values are subjected in a known way to image reconstruction, i.e., the reconstruction of tomograms of slices of, the tomograms being used, for example, for diagnosis.
The detector of the X-ray beam detector system generally contains a number of detector channels, e.g. 1024 detector channels. The detector channels thereby often have non-homogeneities with respect to their radiation sensitivity for X-ray radiation, resulting artefacts known as ring artefacts in the reconstructed tomograms, which are disturbing in the examination and evaluation of the tomograms.
For the suppression of these ring artefacts in the reconstructed tomograms, it is known to use a method in the form of an image post-processing. This method is called an adaptive ring suppression filter (ARU filter), and essentially comprises the following method steps for the post-processing of a tomogram:
a) Screening of bones and air out of the reconstructed tomogram, by limiting all pixel values of the tomogram matrix (designated INB in the following) to a predeterminable upper and lower threshold. This procedural step yields a new image matrix, designated KLB.
b) A first median filtering of the overall image matrix KLB along a multiplicity of straight lines running through the rotational center--in the slice plane--of the computed tomography apparatus, these lines covering the overall image matrix KLB in such a way that a straight line proceeds in the direction of the rotational center from each pixel of the edges of the image matrix KLB. The rotational center need not be located in the tomogram, the tomogram being only the portion of the slice plane shown in the image. The median filtering takes place on the basis of support points that can be predetermined on the straight lines according to image-specific parameters, which respectively represent a pixel value of the image matrix KLB. The pixel values obtained in this way are subtracted from the pixel values of the image matrix KLB, and the difference is subjected to a threshold value evaluation with a predeterminable threshold. The result of this procedural step is a new difference image matrix, designated UDB1.
c) In the region of the rotational center of the reconstructed tomogram, a second median filtering of a submatrix ensues--close to the rotational center--of the image matrices KLB and UDB1 along a number of straight lines proceeding through the rotational center of the computed tomography apparatus, which cover the overall submatrix in such a way that a straight line proceeds in the direction of the rotational center from each pixel of the edges of the submatrix. The median filtering again takes place on the basis of support points that can be predetermined on the straight lines according to image-specific parameters, which respectively represent a pixel value of the submatrix. The pixel values obtained in this way are subtracted from the pixel values of the image matrix KLB, and the difference is subjected to a threshold value evaluation with a predeterminable threshold. The sub-image matrix thus obtained, designated UDB2, is copied into the difference image matrix UDB1, resulting in the difference image matrix designated UDB.
d) Averaging of the pixel values of the difference image matrix UDB on circular arc segments proceeding in both directions from each pixel of the difference image matrix UDB, around the rotational center of the computed tomography apparatus with an opening angle SPHI/2, whereby the opening angle, designated SPHI, of the circular arc segments can be predetermined. The pixels are thereby grouped, according to their distance from the rotational center, into pixels in the inner region and pixels in the outer region of the difference image matrix UDB, in order to enable realization of different opening angles SPHI for the averaging in the inner and outer region.
e) Production of a resultant image in which the averaged pixel values of the difference image matrix UDB, which result from procedural step d), are subtracted from the tomogram matrix INB, and the pixel values obtained are limited to a value range [0,PIXMAX]. The resultant image matrix obtained in this way, designated OUTB, represents the post-processed tomogram with suppressed ring artefacts.
The ARU filter specified above thus in general subjects a reconstructed tomogram to two median filterings and an averaging, but the second median filtering can be omitted if the submatrix, close to the rotational center, of the second median filtering is located outside the represented tomogram. In addition, it is not necessary to median-filter and average the entire tomogram; rather, it is also possible to carry out median filtering and averaging only for regions of the tomogram.
A median filter as a special case of a value of range ordered filter for the suppression of impulse-type disturbances, as defined in lwainsky, Alfred and Wilhelmi, Wolfgang: Lexikon der Computergraphik und Bildnachverarbeitung, Braunschweig/Wiesbaden Friedr. Vieweg & Sohn Verlagsgesellschaft mbH, 1994, p. 235-, ISBN 3-528-05342-9.
The computing steps of the ARU filter, in particular those for median filtering and averaging, are carried out successively in line-by-line fashion for each pixel of the tomogram, without intermediate results of calculations of the neighboring pixels being usable. FIGS. 1 and 2 herein illustrate the line-by-line sequential procedure in the execution of the median filtering and the averaging of an image matrix (in the Cartesian grid) of a tomogram. On the basis of this procedure, in the calculations associated with the median filtering and the averaging, a high number of memory accesses to (retrievals of) the pixel values (in the Cartesian grid) of the tomogram are required, as are a high number of computing operations. In the computing steps of the ARU filter, the index determination of the required input data (pixel data) takes place in floating-point fashion, and the access to the input data (pixel values) located in the discrete Cartesian grid takes place by means of nearest-neighbor interpolation of the support points at the corresponding pixels. In the averaging on circular arcs, in order to conserve or limit computing time it is thus necessary to limit the averaging to a relatively small number of pixels, however, this has a detrimental effect on the quality of the suppression of the ring artefacts in the tomogram.
An improvement of the above-described method is to reinterpolate the input data of the ARU filter, which are present in the form of an image matrix of the tomogram in a Cartesian grid, into a polar grid with a constant increment of angle and radius. In the ARU filter, this enables a simple sequential access to the input data, since in the ARU filter filtering takes place in the radial direction and averaging takes place in the azimuthal direction (cf. FIGS. 3 and 4).
By means of the sequential memory access, the pixel values of the post-processed tomogram in the polar grid can now be calculated iteratively from the results of previously calculated pixel values. The intermediate results of calculations of the adjacent pixels are usable, so that a sharp reduction of the computing operations to be carried out and of the memory accesses to the input data in the calculation of a pixel value results. At the end of the calculation of the ARU filter, the output data are again reinterpolated onto a Cartesian grid before computing step e).
The disadvantage of this version of the method with the reinterpolation onto a polar grid with a constant increment of angle and radius is, however, the resolution of the polar grid that is required so that the original resolution of the input data is not lost in the reinterpolation from the Cartesian grid. In order to avoid having to accept any loss of resolution of the input data, the polar grid has to be defined in such a way that the constant increment of angle and radius at the tomogram pixel that is furthest away from the rotational center corresponds to the local resolution of the Cartesian grid. This has the consequence that the data set of input data has to be increased by an order of magnitude in the reinterpolation from the Cartesian grid to the polar grid, since image regions located closer to the rotational center are then scanned more finely than they were in the Cartesian grid. The consequence of the increase in the data set, required for reasons of resolution, is an increase of computing operations due to the increase in the number of pixels. This, however, counteracts the reduction in the computing operations achieved by means of the iterative calculations in the calculation of a pixel.